There are many devices which now use radiation detectors. For example, CT scanning has proven invaluable for medical diagnosis and analysis and is now in wide use. However, the newest CT scanning techniques tend to be somewhat expensive, due in part to the cost of the components normally associated with a large number of radiation detectors.
CT scanners are utilized to provide a computed cross-sectional detail of soft living tissue structures. Briefly, the cross-section of interest is positioned between a radiation source (e.g., an X-ray tube) and a detector system. A portion of the beam is absorbed by the tissue during transit along ray paths through the section. Thus, absorption by the body section along any given path is a function of the sum of the absorption coefficients of the particular body tissues through which the beam passes. That portion of the radiation which passes through the section is detected by, for example, a scintillation crystal which produces light photons in response to incident radiation. Typically, the scintillator is optically coupled to a photomultiplier tube (PMT) which converts the light photons into electrical output signals. Such measurements are made along many paths through the body section to provide data used in calculating an array of point-by-point relative absorption coefficients. The computed coefficients are then utilized to provide a visual display of the cross-section.
Some of the earlier CT scanners obtain the requisite multiplicity of absorption measurements by synchronously scanning the body section with a highly collimated X-ray beam and an aligned single detector. The beam and detector are together translated to scan the body section and generate a set of measurements along parallel paths. This assembly is then rotated with respect to the body section and the translation-rotation operation is repeated to provide sets of measurements at different angular dispositions.
It is desirable, however, to minimize the overall scanning time. Accordingly, more recent CT scanners generally utilize a plurality of individual radiation detectors in co-operation with a fan-shaped X-ray beam wide enough to irradiate the entire body section. In one example of such systems, a fan-beam source rotates about the body section, irradiates a stationary detector array forming the outside perimeter of the scanning frame. As many as 600 scintillation crystals and associated photomultipliers tubes (PMT) are typically used in such systems. Thus, the cost of the scintillation crystal-PMT detectors, and particularly the photomultiplier tubes, presently represent a substantial portion of the cost of such CT scanners. Accordingly, it is desirable to directly convert the radiation to an electrical signal and eliminate the need for large numbers of photomultiplier tubes and associated circuits.
Various direct conversion detectors such as Xenon gas and high purity germanium (HPGe) semiconductors have been utilized in the past. However, such detectors are disadvantageous for other reasons.
The properties of high purity germanium are well known. A high density position sensitive array has been described in "Two Detector 512-Element High Purity Germanium Camera Prototype" Kaufman et al, IEEE Transactions Nuclear Science, NS-25, February, 1978. (Presently in press) However, the production of high purity germanium is a relatively complex and costly procedure, and further high purity germanium detectors require special cooling apparatus, (e.g. liquid nitrogen). Additionally, it is desirable in CT scanning applications to use a thinner detector than generally can be made with high purity germanium.
Room temperature semiconductor detectors would thus appear to be a natural choice for application in CT scanners. Since photon energy from the radiation is directly converted into an electrical charge, compact arrays could be assembled without the requirement of more bulky photomultiplier tubes. Such higher detection element density is especially advantageous for CT scanning applications. These high density semiconductor detector arrays could also be mass produced by batch processing.
However, semiconductor detector materials now available in reasonable quantities have been considered by others in the past and discounted as unsuitable for CT scanning applications. For example, cadmium telluride (CdTe) has specifically been studied with respect to use as a detector in nuclear medicine, and has been found not compatible with the needed accuracies. Specific reference is made to Allemand et al, "Present Limitations of CdTe Detectors in Nuclear Medicine", Revue de Physique Appliquee, 12: 365-367, February, 1977. Other semiconductor detectors are discussed in Armantrout et al, "What Can Be Expected from High-Z Semiconductor Detectors", IEEE Transactions on Nuclear Science, Vol. NS-24, No. 1, Feb. 1977. Of the semiconductor materials that are described, various ones have been tested and show unsuitable characteristics similar to CdTe. Computer studies project that similar properties are present in other semiconductor detectors--not presently tested.
More specifically, CdTe detectors have been found to exhibit adverse "leakage current", "Tailing", "polarization" and "memory" properties, which are generally considered to render CdTe unsuitable for CT scanning usages. Other detectors, such as H.sub.g I.sub.2, have been found to exhibit similar properties.
The high "leakage current" property varies with temperature, and further, varies with time once bias has been applied to the detector. The "tailing" phenomenon causes the detector response to portray a monochromatic (single energy) source as having energies ranging all the way down to noise levels. "Tailing" increases as a function of the elapsed time after bias is applied to the detector. The "polarization" property is manifested by the most frequently encountered detector response (hereinafter termed the photo-peak) to a constant monochromatic source, shifting towards the low energy end of the spectrum. The "memory" property is manifested by a continued current flow after a pulse of incident radiation has ceased, sometimes with a time constant of hours.
Various techniques have been proposed for dealing with such adverse properties. However, the proposed techniques have tended to merely add to the complexity of the detector circuitry and, in fact, to introduce sources of errors that have made application of CdTe detectors in CT scanners impractical. For example, the high current leakage could be compensated by sampling the detector output just before the X-ray beam is activated to produce each pulse of incident radiation. However, such a scheme will compromise the signal-to-noise ratio of the system when absorption is high, as happens when the beam goes through bone. Pre and post measurement calibration techniques for memory and polarization compensation, similarly tend to add signal-to-noise problems. Further, polarization compensation techniques wherein the bias to the detectors is turned off between measurements, are inadequate due to hysteresis effects.